The present invention is related to methods and apparatus for surgically implantable pumps that provide a mechanical device for augmenting or replacing the blood pumping action of damaged or diseased hearts. More specifically, this invention is related to methods and apparatus for conduits for such pumps to meet the three-fold requirements of a substantially non-delaminable conduit having a highly non-thrombogenic porous inner surface that is not susceptible to air leakage under negative pressure, whereby the use of such implantable pump procedures is broadly enabled, and wherein the functional utility, ease of use, and wide applicability of the device in medical practice constitutes progress in science and the useful arts. Furthermore, the present invention teaches processes to make and use the device in medical practice.
Of all the cardiovascular disorders, congestive heart failure (CHF) is the only one to show a sharp increase in prevalence since the 1960s. This rise in the number of cases of CHF worldwide is a major and growing public health concern. While the prevalence of coronary heart disease (CHD) has declined in the past few decades and the prevalence of stroke has remained steady until recent years, the number of people suffering from CHF has increased dramatically. In terms of mortality and morbidity, the prognosis for CHF is depressingly poor and in the US, the disease has been recognized as ‘epidemic’. The high costs associated with the condition will place an added burden on public health resources as the incidence of CHF continues to rise.
By definition, CHF is a disorder in which the heart fails to pump blood adequately to other organs in the body. This can result in a shortness of breath, fatigue and fluid retention (edema) and if left unchecked can lead to death within a few years. CHF is not a disease per se but a condition that arises as a result of various cardiovascular diseases (CVD). In effect, CHF is the end-stage syndrome of heart muscle disorders and diseases such as hypertension and CHD, which can damage or impair the functionality of the heart and the vessels supplying it.
In the majority of cases, CHF is a progressive condition and over time, the ability of the heart muscle to function properly deteriorates—this process is called cardiac remodeling. As the condition worsens, the ventricular muscle over-stretches and the muscle fails to work to its full efficiency. This leads to a further reduction in the cardiac output and exacerbates symptoms of heart failure.
In CHF, the reduced cardiac output causes a fall in arterial pressure leading to the activation of several compensatory reflexes. The sympathetic nervous system is stimulated, resulting in a direct increase in the force of contraction of the heart and a greater venous return as a response to venoconstriction. Long-term compensation includes the activation of the renin angiotensin system (RAS) and subsequent renal fluid retention. The combined effect of these responses can lead to the formation of edema, especially in the legs and ankles. If heart failure occurs in the left side of the heart, pulmonary edema can result which manifests as breathlessness. In advanced CHF, the severity of the symptoms can be disabling and often leads to hospitalization. An added consideration is sudden cardiac death, which can occur at any time during the course of CHF.
The most common classification of CHF is based on criteria set out by the New York Heart Association (NYHA). Originally published in 1928, the classifications defined the stages of CHF by its clinical severity and the functional status of the cardiac muscle. Over the years the classifications have been updated and the latest revision was in 1994. The NYHA classes of CHF are listed below.
New York Heart Association Functional Classifications of CHF
Class I Patients with cardiac disease but no resulting limitation of physical activity. No fatigue, palpitations, shortness of breath or angina during normal physical activity.
Class II Patients with cardiac disease with slight limitation of physical activity. Normal physical activity results in symptoms of heart failure or angina but the patient is comfortable at rest.
Class III Patients with cardiac disease with marked limitation of physical activity. Slight physical activity results in fatigue, palpitations, shortness of breath or angina but comfortable at rest.
Class IV Patients with cardiac disease and an inability to carry out physical activity without discomfort. Even at rest, the symptoms of heart failure or angina may be present.
The failing heart is a result of a number of factors combining to reduce the efficiency of the heart as a pump. The most common dysfunction is an impairment of left ventricular function, which is present in 80-90% of patients with CHF. As the blood flow from the heart slows, the blood returning to the heart through the veins can back-up, resulting in congestion in the tissues. This can lead to swelling in the legs and ankle and fluid retention in the lungs, which interferes with breathing and contributes to the characteristic shortness of breath seen in people with CHF.
Several types of surgically implantable pumps have been developed in an effort to provide a mechanical device for augmenting or replacing the blood pumping action of damaged or diseased hearts. Some of these pumps are designed to support single ventricular function. Such pumps usually support the left ventricle, which pumps blood to the entire body except the lungs, since it becomes diseased far more commonly than the right ventricle, which pumps blood only to the lungs. Other devices have been tested and used for providing biventricular function.
Depending on the needs of a particular patient and the design of a pump, pumping units such as so-called “VADs” (ventricular assist devices) can be implanted to assist a functioning heart that does not have adequate pumping capability. Left-ventricular assist devices (LVAD) in particular are recognized as potentially very valuable for assisting patients who suffer from congestive heart failure. An LVAD is able to fully take over the function of the left ventricle, thus perfusing the body with oxygen-rich blood.
The LVAD attaches to the patient's natural heart, and to a natural artery, and can be removed if the natural heart recovers. Some LVADs are surgically implanted into the patient's abdominal cavity, while others remain outside the body and are placed in fluid communication with the heart via elongated cannulas. Recently, a National Institutes of Health study estimated that as many as thirty-five thousand people could be candidates for use of a left-ventricular assist device.
At present, conventional ventricular assist devices are used for patients who are waiting for a heart transplant (a so-called, “bridge to transplant”), or alternatively to patients whose natural heart is of such poor condition that the patient cannot be removed from a heart-lung machine without providing some assistance to the patient's heart following otherwise successful open-heart surgery. Still another group of patients eligible for the use of conventional ventricular assist devices are those who suffer massive heart attacks that lead to circulatory collapse. The suitability of long-term utilization of conventional left-ventricular assist devices outside of the clinical environment remains under study.
Expansion and contraction of a variable-volume chamber typically effect blood flow in the LVAD. One-way valves associated with the inflow and outflow ports of the LVAD permit blood flow propelled by the natural left ventricle into the variable-volume chamber during expansion, and blood flow out of this chamber, usually to the ascending thoracic aorta. These one-way flow valves may be constructed as part of the LVAD itself, or may be disposed in separate blood-flow conduits attached thereto. A pair of artificial blood conduits respectively connect the inlet port of the variable-volume chamber (or the inlet end of a valved conduit) to the left ventricle and the outlet port of the variable-volume chamber (or the outlet end of a second valved conduit) to the major artery which is to receive the blood flow from the device.
As is well known, artificial blood conduits have become a valuable tool of modern medicine. One use of such artificial blood conduits is as a temporary or permanent prosthetic artery. Another use is in the connection of temporary blood pumps, such as ventricular assist devices described herein, between the left ventricle of the heart and a major artery.
The demands on artificial blood conduits in ventricular assist devices are great. The conduit must deal with the pulsatile blood flow created by the host's own heart, as well as with the flow, pressure, and pulsations created by the assist device. Moreover, there are differences in flow and pressure between the inflow and outflow conduits connected to the pumping device. For example, while the outflow conduit experiences regular pulses of high pressure, flow in the inflow conduit is dependent on the pumping strength and rhythm of the natural left ventricle on top of which the periodic LVAD pressures are superimposed (i.e., expansion of the variable volume chamber tends to pull fluid from the inflow conduit). The inflow conduit thus sees irregular and typically low flows and pressures; here, the negative pressure transients that can occur in the inflow conduit are of special importance.
Conventional artificial conduits for use in LVADs may be constructed of an elongate flexible woven polyethylene terephthalate (PET) fabric tube. In some cases, the conduits are sealed with a thin bio-compatible collagen coating on the inner lumen wall to render the fabric more leak resistant at the time of implantation, and also more compatible with the patient's blood. The collagen coating, typically bovine collagen, eventually is absorbed into the blood stream and is replaced with a natural coating of blood cells, serum protein, and other elements from the blood. In the absence of a sealant, the conduit may have to be pre-clotted by the surgeon just prior to implantation.
As is generally known in the art, a porous surface on the inner lumen wall of an implanted blood conduit is advantageous because it becomes coated with the natural coating of blood cells, serum protein, and other elements from the blood. This coating inhibits clot formation (thrombogenesis) which is highly desirable. Earlier inflow conduits comprised PET woven grafts surrounded by a machined flexible housing. However, such conduits alternately collapsed and regained their tubular shape with each stroke of the pump. Prior art improvements comprised creating a scaffolding which prevented the collapse of the conduit with each pump stroke. however, since the tube no longer collapsed, the lumen of the tube experienced negative pressure and atmospheric air consequently infiltrated through the walls into the lumen.
Porosity also renders an inflow conduit vulnerable to the entrance of air during the intervals of negative pressure transients that can occur in the inflow conduit as noted above. This raises a serious problem whenever the inflow conduit is in contact with the atmosphere. For example, the surgical implantation of the device entails operation of the LVAD or VAD during the period when the chest of the patient is open to the atmosphere. Likewise, externally mounted LVAD's may have inflow conduits in contact with the atmosphere. Air can be sucked into the inflow conduit during intervals of negative pressure transients as outlined above, leading to air embolisms in the cardiovascular system of the patient undergoing the implantation procedure. Such air embolisms can lead to injury to the patient or even to the death of the patient. It is obvious, therefore, that this risk of air embolisms resulting from the vulnerability of the inflow conduit to the entrance of air during intervals of negative pressure, whenever the inflow conduit is in contact with the air, is an extremely serious problem. Even if the conduits are sealed with a thin bio-compatible collagen coating on the inner lumen wall to render the fabric more leak resistant at the time of implantation this procedure only addresses the problem of leakage of the lumen contents outward, and does not at all address the problem of the entrance of air during the intervals of negative pressure transients that can occur in the inflow conduit as noted above.
Some non-implantable ventricular assist devices utilize cannula-like conduits that are relatively rigid, some being formed of smooth, reinforced non-porous polyurethane. Such conduits might solve the problem of leakage into the conduit lumen when it is under negative pressure, but they would not be suitable for use in implantable devices, as they will not easily accommodate varying anatomical placements, and tend to kink if bent.
My prior invention of an implantable ventricular assist device disclosed in U.S. Pat. No. 6,001,056, which is expressly incorporated herein in its entirety by reference, comprises an inflow conduit. The inflow conduit includes a flexible tubular graft body having an upstream end and a downstream end, the body having a substantially smooth inner surface for enhanced flow-through of blood with a minimum of surface-induced turbulence. The inflow conduit also includes a ventricular attachment structure to which the upstream end of the body connects, and a coupling fitting on the downstream end of the body. An implantable pumping portion may be placed in flow communication with the inflow conduit and with an outflow conduit. The tubular graft body may be a knitted fabric having a biocompatible sealant impregnated therein, or a closed structured PTFE.
Even in the case of an inflow conduit fabricated from closed structured PTFE in which the tubular wall of the conduit has a pore size of not less than 2.mu., the water entry pressure for the base tube is still at least about 5 psi. Even when a thin PTFE tape having a thickness of about 0.01 mm and an ethanol bubble point of at least about 2 psi is wrapped about and laminated to the base tube the resulting extremely low porosity tubular wall still leaks air into the lumen of the tube when placed under negative pressure. Thus, even this device does not solve the twin requirements of a conduit having a highly non-thrombogenic porous inner surface that is not susceptible to leakage under negative pressure.
My prior invention of a non-porous smooth ventricular assist device conduit solves this problem. It is disclosed in copending U.S. patent application Ser. No. 09/874,846 filed Jun. 5, 2001, which is expressly incorporated herein in its entirety by reference, is directed to an improved inflow conduit for an implantable ventricular assist device having a flexible, porous tubular graft body that has an upstream end and a downstream end. The tubular graft body has a substantially smooth inner surface for enhanced flow-through of blood with a minimum of surface-induced turbulence, a ventricular attachment structure to which the upstream end of the body connects, and a coupling fitting on the downstream end of the body. The improvement comprises covering the flexible porous tubular graft body with an attached non-porous polymer, whereby a nonporous conduit is formed. The attached non-porous polymer may be attached by thermal bonding, or by a biocompatible adhesive.
Although this latter conduit uniquely solves the problem of air leakage, the possibility of delamination of the attached non-porous polymer of the conduit of the invention could once again permit such leakage. In view of the absence of methods and apparatus for conduits that meet the three-fold requirements of a substantially non-delaminable conduit having a highly non-thrombogenic porous inner surface that is not susceptible to air leakage under negative pressure, there remains a need for a substantially non-delaminable smooth ventricular assist device conduit and process for making same.
Thus, in spite of extended efforts in academic medicine and the pharmaceutical industry, there remains room for improvement in the construction and function of conduits for ventricular assist devices. Even though conduits for LVAD's are used extensively in medical practice, prior devices, products, or methods available to medical practitioners have not adequately addressed the need for non-delaminable apparatus for conduits for LVAD's to meet the need for methods and apparatus for conduits for such pumps as outlined above. The present invention embraces and finally addresses the clear need for advanced non-delaminable apparatus for conduits for LVAD's and VAD's to meet the three-fold requirements of a substantially non-delaminable conduit having a highly non-thrombogenic porous inner surface that is not susceptible to air leakage under negative pressure as set forth above. Thus, as pioneers and innovators attempt to make methods and apparatus for LVAD's safer, cheaper, more universally used, and of higher quality, none has approached same in combination with simplicity and reliability of operation, until the teachings of the present invention. It is respectfully submitted that other references merely define the state of the art or show the type of systems that have been used to alternately address those issues ameliorated by the teachings of the present invention. Accordingly, further discussions of these references has been omitted at this time due to the fact that they are readily distinguishable from the instant teachings to one of skill in the art.